The goal of radiotherapy is to irradiate unhealthy, usually cancerous, tissue or tumors with the aim of killing that tissue while sparing, to the extent possible, the surrounding healthy tissue. This requires that the patient be irradiated with one or more radiation beams that are carefully positioned and shaped in order to precisely deliver the intended quantity of radiation to a specific target volume. In dimensions transverse to the nominal direction of the beam, this shaping is often performed with metal apertures or multi-leaf collimators that define the profile of the beam. The beams are often directed from different angles at the patient, thereby increasing the radiation dose in the intersecting regions. A treatment plan is generated for each irradiation session of a patient that describes the various beam angles, shapes, energies, and quantity of radiation delivered along with the associated maps showing the radiation dose intended for the patient.
With high-energy X-ray beams, the technique of spatially shaping the radiation dose with multi-leaf collimators and multiple beams directed from multiple angles is known as Intensity Modulated Radiation Therapy or IMRT.
Charged particles, typically protons but also heavier ions, are also used for delivering spatially complex radiotherapy treatments. The particle accelerators typically produce a narrow, or pencil, beam that can be focused, collimated, or deflected with magnetic optics, although sometimes the beam is broadened in size or angle by scattering the pencil beam. With particle beams, the shaping can also be effected along the beam direction by varying the energy distribution of the charged particles, as the penetration distance and dose distribution along the particles' tracks depends on their energy. For a monoenergetic ion beam in a homogeneous medium, the absorbed dose versus depth describes a curve known as a pristine Bragg peak as seen at 12 in FIG. 1.
From entrance into the body until close to the penetration depth, the delivered dose, which is proportional to the average linear energy transfer (LET) given in keV/μm due to ionization by the beam, rises slowly along the plateau of the curve. Near the final depth, the LET increases rapidly to a peak, then falls to zero rapidly at the end of the track of the ensemble of particles in the beam, thereby increasing the radiation dose in a narrow volume. The LET value is typically a characteristic of a beam as a function of depth. It is related to the energy loss per unit length dE/dx per unit density, typically given in MeV cm2/g, of each charged particle. dE/dx is a function of the particle energy.
In this specification, the term beam generally refers to a penetrating radiation beam. However, it may either refer to a radiation beam produced by an accelerator, a narrow pencil beam at the output of the beam delivery system, a set of pencil beams delivered within a short time, i.e., a composite beam, or a broad beam restricted by a metal aperture or collimator. A treatment field generally refers to a composite or broad beam with a set of particle energies and a defined spatial distribution delivered within a short time. If either the energies or spatial distribution or both are changed, that generally refers to a new treatment field.
By irradiating with beams having different energies, or by modulating the beam energy as it is applied, the dose distribution can be shaped to deposit maximal dose in an extended volume containing the tumor while delivering reduced dose to the surrounding healthy tissue. An example, shown at 14 in FIG. 1, uses a beam with a set of different energies and fluences arranged to produce a depth-dose curve with a flat maximum, denoted as a spread-out Bragg peak or SOBP. The energy of particles in the beam can be reduced by inserting absorbing elements, typically plastic, upstream of the patient. The energy reduction depends on the thickness of the absorber, so a set or range of energies can be produced with a spatially varying thickness to the absorber or by rotating a wheel with multiple thickness steps and temporally modulating the beam energy.
FIG. 1 highlights one of the differences between radiotherapy using x-rays or charged particles. For an equivalent dose at the depth in the middle of the SOBP, the x-ray beam deposits more dose in front of that depth and after the tail of the Bragg peak, represented by the area between the x-ray dose curve 10 and the charged particle SOBP dose curve 14. This ability to limit the dose to healthy tissue outside of a tumor is a major motivation for delivering radiotherapy with charged particles.
Charged particle beams can also be scanned across the patient if the beam is small enough and scanning magnets are provided to steer the beam in angles away from the central axis of the system. This technique is generally termed pencil beam scanning or PBS. PBS treatments can be delivered without metal apertures or collimators if the scanning magnets can sufficiently define the transverse shape of the composite beam. Typically, more than one radiation field will be applied from a single direction with different energies or energy spectra used for the different fields, thereby shaping the delivered dose at different depths.
Several different forms of PBS are used to deliver radiotherapy. A technique known as Intensity Modulated Proton Therapy or IMPT allows the beam intensity to vary as a function of lateral position by modulating the beam current or modulating the scanning rate. In some scenarios, the beam position is continually scanned and in others it is translated to a series of fixed positions where the beam is turned on for a period of time, a technique called spot scanning. For this approach, the magnetic scanning of the beam defines the shape of the irradiated volume transverse to the beam direction. An approach, which is termed either wobbling or uniform scanning, has the beam scan a given area, typically a regular region like a circle or rectangle, and deliver a uniform dose over that area. A fixed aperture or multi-leaf collimator is positioned between the beam source and the patient to restrict the transverse distribution of radiation.
In all of these various forms of radiotherapy, the standard for positioning and shaping the beam relative to the patient is typically less than one millimeter across a field of regard that can range up to 40 cm in diameter. This requires that the radiation pattern be delivered with precise shaping and positioning of the beam(s) over those scales, as well as in precise amounts so the dose absorbed by the patient matches the prescription of the treatment plan. This requires careful calibration and measurement of the delivery system and the treatment plan, i.e., the beam(s) which will be delivered to the patient, a process referred to as quality assurance or QA. QA measurements are routinely performed that reassure the practitioners that the correct radiation dose will be delivered in the correct amount over the correct spatial distribution to the patient. These measurements are performed with a variety of different radiation detectors.
With the complex series of radiation fields that are used in radiotherapy, especially in charged particle therapy, it is critical to make measurements which predict the absorbed dose inside of a patient, not merely in a plane just outside of the patient. One way this is accomplished is by placing absorbing material, a tissue phantom, in between the beam source and the detector so the radiation incident on the detector corresponds to the radiation incident on the corresponding location(s) inside of the patient. The absorbing material is typically water (since the body is mostly made up of water), sheets of plastic with water-equivalent absorbance, or more complex phantoms with multiple materials.
There exist several different detectors used for performing QA of these radiotherapy beams. Film has been used since the beginnings of radiotherapy to view the beam as delivered to the patient. Various other technologies have been used, including scintillating screen based portal imagers, flat panel devices consisting of a sheet of scintillating crystals atop an array of semiconductor detectors, ionization chambers as single devices, stacked along the beam direction, or arrayed transverse to the beam direction, multi-wire proportional chambers, gaseous electron multiplier (GEM) detectors, and other detectors, etc. Typically, these detectors are one part of the overall system and procedure for performing QA.
Requirements for these detectors include stability, robustness, linearity, accuracy, and precision, i.e., signal to noise ratio (SNR), spatial resolution, and dosimetric accuracy. The detector output should be translatable into a measure of the effective dose delivered to the patient. In addition, marketability factors such as the cost and usability of the detector output are important for the usefulness of a detector for QA applications.
An ideal QA detector would measure the delivered dose inside a tissue phantom with sufficient spatial resolution to validate the treatment plan. In practice, current detectors fall short in one or more respects. Some of the detectors employed, like film and detectors using thin scintillating screens or sheets of crystal, have spatial resolution that exceeds the requirement for beam positioning and shaping. However, their detection signal is not linear with the fluence or LET of the radiation beam, which is required in order to predict the corresponding dose to the patient.
The light output of some scintillators will saturate when a clinically useful fluence is incident. Inorganic scintillators or phosphors such as P11 (ZnS:Ag), P20 (Zn,CdS:Ag), P43 (Gd2O2S:Tb), P46 (Y3Al5O12:Ce), and P47 (Y2SiO5:Ce) have responses that are typically linear with fluence. However, scintillator light output is generally not linear with the LET of the incident particles. A plot of the light output versus depth has the same general shape as the Bragg peak. However, the ratio of the peak value to the value at a point on the plateau differs from the ratio of LET values calculated using accurate physical models or that is measured by a standard detector such as an ion chamber. When the ratio of light output is less than the ratio of LET or absorbed dose, this is termed quenching. The physical explanation is that an abundance of excited atoms provide alternate pathways for excited species to relax by means other than photon emission.
Scintillating screens are used, in conjunction with radiation therapy, for absorbed dose measurement (as described, for example, by J. M. Schippers, S. N. Boon and P. van Luijk, “Applications in Radiation therapy of a scintillating screen viewed by a CCD camera,” Nucl Instr. and Meth. A 477, pp. 480-85 (2002), and S. N. Boon, thesis, “Dosimetry and quality control of scanning proton beams” (1998), available online at http://www.ub.rug.nl/eldoc/dis/science/s.n.boon/, and references therein, all of which are incorporated herein by reference). They are also useful for monitoring beam delivery in real time (S. M. Ebstein, High Resolution Proton Beam Monitor, U.S. Pat. No. 7,515,681). However, as those references show, knowledge of the particle beam energy spectrum and careful calibration of the detector response is required to make accurate dosimetric measurements.
Other detectors such as ion chambers are quite linear in their measurement of the fluence and LET, so they are quite accurate in predicting absorbed dose. However, due to their size and complexity, they cannot be made into arrays that can measure the incident radiation distribution with the required resolution.
Due to the limitations of these detectors, a complex system is required to perform and evaluate the QA measurements. Typically, a series of representative measurements are made with detectors that have high spatial resolution but poor dosimetric accuracy or detectors with limited spatial resolution but good dosimetric accuracy. The response of the detectors to the delivered radiation beams is calculated from a model of the beam delivery and detection system and compared to the actual measurements. This approach is described in C. Brusasco and B. Marchand, Device And Method For Particle Therapy Monitoring And Verification, U.S. Patent Application Pub. No. US 201110248188 A1.
However, this approach cannot give complete confidence that the delivered beam matches the plan due to reduced dimensionality of the measurements. In addition, the work required to construct the models and perform the calculations of the expected measurements adds to the cost and complexity of performing QA.
There exists a need for improved detectors, especially for charged particle therapy QA. The beam delivery systems are complex instruments and are required to produce a series of beams with complex shapes and different energy spectra. Current detectors and systems cannot directly measure the delivered dose as a function of depth with sufficient accuracy and spatial resolution to quickly and easily perform the QA task. A detector which combines the high spatial resolution of either film or a scintillating screen detector and the dosimetric accuracy of an ion chamber would be very useful in performing QA of complex radiotherapy treatment plans, especially for IMPT.